Heart disease and stroke, which are the principal components of cardiovascular disease, remain the leading cause of death in the western world. One of the most common treatments for coronary artery disease is coronary bypass graft surgery (CABG), where a suitable length of the patient's saphenous vein or the internal thoracic or mammary arteries are used to supply blood to the heart tissue. The number of CABG procedures in the US was more than 600,000 (1.2 million worldwide) in the year 2000, but these tissue grafts tend to deteriorate due to further advancement of the patient's coronary artery disease and disruption of the normal vascularity.[1-3] On the other hand, total peripheral artery bypass grafting is performed to relieve the symptoms of vascular deficiencies, where a common problem involves the supply of autologous bypasses. The lack of nondiseased saphenous veins as arterio-venous access fistulae for haemodialysis is a major cause of morbidity for patients with renal failure.[4, 5]
From the composition point of view, it is important to note that cardiovascular tissues are composite materials with elastin and collagen as the main load bearing components. The directional manner in which collagen fibers are arranged within the tissue creates the anisotropic behavior, with higher tensile strength in the circumferential than in the axial direction.[6-9]. Accordingly, the mechanical properties of soft tissues, including aortic tissue, are anisotropic, with a higher stiffness in the circumferential than in the axial direction.
For a biomaterial to be used as tissue replacement, it is important to ensure a good match of the mechanical properties of the implanted device and the surrounding tissues [10]. Elastic polymers have been investigated to create compliant grafts since the mismatch of the native aorta and the synthetic grafts, such as Dacron and ePTFE, may contribute to intimal hyperplasia (IH) and ultimate failure. A successful replacement has been reported as having adequate strength, kink resistance, and must allow sutures to hold under circumferential and axial tension, as well as circumferential and axial compliance. Difference in compliance results in haemodynamic changes and increased shear stresses that may induce the release of growth factors that stimulate IH.[10-12].
Even though there are several FDA approved materials for replacement aorta, such as Dacron or e-PTFE, these materials do not posses the same tensile properties as the tissue they are replacing, which results in hemodynamic problems and mismatch of mechanical properties and other problems at the implant/tissue junction. Therefore, it would be very advantageous to be able to produce a material that provides a suitable match with the mechanical and viscoelastic properties of biological tissues. Two promising material systems for meeting this need are anisotropic hyrdrogels and nanocomposite hydrogels, as described below.
Hydrogels
Hydrogels have been shown to be promising candidates for a wide range of biocompatible tissue replacement materials. Hydrogels are hydrophilic polymer networks produced from reactions of one or more monomers or by association bonds between chains that can absorb from at least 20% to up to thousands of times their dry weight in water [13, 14]. Hydrogels may be chemically stable or they may disintegrate and dissolve with time. They are called either physical (reversible) or chemical (permanent) hydrogels. Physical hydrogels have networks held together by molecular entanglements and/or secondary forces such as hydrogen bonding, van der Waals interactions, ionic or hydrophobic forces. Physical hydrogels are not homogeneous due to regions of high crosslinking density and low water swelling, called clusters, dispersed within low crosslinking density and high water swelling, or hydrophobic or ionic domains that create inhomogeneities. Chemical hydrogels are covalently crosslinked networks, but they may also be generated by crosslinking of water-soluble polymers, or by converting hydrophobic polymers to hydrophilic polymers. Chemical hydrogels are also not homogeneous due to clusters of molecular entanglements. Chain loops and free chain ends also produce network defects in both physical and chemical hydrogels, and they do not contribute to the permanent network elasticity [13, 15].
The main areas in which hydrogels are used as biomaterials is in contact lenses, synthetic wound coverings, drug delivery systems, organ and tissue replacements, and permselective membranes [13, 16, 15, 17-24].
An important characteristic of hydrogels is their swelling behaviour in water, since after preparation they have to be in contact with water to yield the final solvated network structure. Highly swollen hydrogels are those of polyvinyl alcohol (PVA), polyethylene glycol, and poly-N-vinyl 2-pyrrolidone, among others. PVA is a hydrophilic polymer with various characteristics desired for biomedical applications, such as high degree of swelling, uncomplicated chemical structure, rubbery/elastic nature, and non-toxic. PVA can be converted into a solid hydrogel by crosslinking.
Crosslinking can be accomplished by using several methods. For biomedical applications, physical crosslinking has the advantages of not leaving residual amounts of the toxic crosslinking agent, and higher mechanical strength than the PVA gels crosslinked by either chemical or irradiative techniques.
The mechanical properties of the PVA hydrogels are similar to that of soft tissue, including elasticity and strength, and can be controlled by changing the number of thermal cycles, PVA concentration, thawing rate of the thermal cycling process, and freezing holding time among other parameters [19, 25, 26]. A PVA based bioprosthetic heart valve stent has been fabricated. However, the mechanical strength and stiffness of these PVA materials were weak and did not fully match the mechanical properties displayed by the cardiovascular tissues such as arteries and heart valves.
PVA has a relatively simple chemical formula with a pendant hydroxyl group and a crystalline nature, which allows it to form a solid hydrogel by the crosslinking of the PVA polymer chains. Vinyl alcohol (monomer) does not exist in a stable form and rearranges to its tautomer, acetaldehyde. PVA is produced by free radical polymerization of vinyl acetate to polyvinyl acetate (PVAc), and subsequent hydrolysis of PVAc gives PVA [25].
PVA can be crosslinked using several methods, such as the use of crosslinking chemical agents, using an electron beam or γ-irradiation, or the physical crosslinking due to crystallite formation. For biomedical applications, physical crosslinking has the advantages of not leaving residual amounts of the toxic crosslinking agent, and higher mechanical strength than the PVA gels crosslinked by either chemical or irradiative techniques [27, 28]. In chemical cross-linking, the chemical agents that react with the hydroxyl groups are glutaraldehyde, ethylaldehyde, terephthalaldehyde, formaldehyde, hydrochloric, boric or maleic acid, among others [19, 29].
Physical crosslinking forms a hydrogel with a network of semi-crystallites of hydrogen bonds of polymer filled with solvent [30]. It has been shown that the mechanical properties of the hydrogels, including elasticity and strength, can be altered by changing the PVA concentration, the number of freeze/thaw cycles, the process thawing rate, the freezing holding time, and the freezing temperature [19, 29, 30]. Increasing the PVA concentration results in hydrogels with higher crystallinity and added stability upon swelling, which increases its tensile strength and tear resistance. The lower the initial concentration of PVA, the fewer the polymer chains in solution, and there may be less number of crystalline regions created in the cycled PVA. Increasing the number of freeze/thaw cycles increases the strength and stiffness of the hydrogel by reinforcing existing crystals within the structure [19, 25, 26]. Decreasing the thawing rate of frozen PVA solutions increases the tensile strength because the solutions are kept for longer periods at temperatures below 0° C., allowing for increasing movements of polymer chains which result in further entanglements and increased crystallite size and numbers.
The freezing holding time also has a drastic effect, with samples frozen up to 10 days giving the most mechanically strong PVA hydrogels [19, 26, 28, 29]. The freezing temperature has an interesting effect. The freezing temperature controls the phase equilibria and dynamics, where the lower the temperature of the system the lower the amount of unfrozen solvent in the liquid regions. Therefore, the lower the temperature the less opportunity for chain mobility in the polymer rich regions, giving less chances of crystallite growth and formation. This explains why keeping the frozen PVA solutions at −10° C. produces somewhat more rigid hydrogels than those kept for the same period of time at −20 or −30° C. The freezing rate was shown not to have drastic effects on the properties of the hydrogel [19, 26, 30]. PVA hydrogels not only have tensile strength and elongation, but also flexibility and elasticity. Research has proven its ability to recover to its original shape after being deformed to strains of 50%, showing excellent persistence and repeatability of the recovery [30].
Physical crosslinking allows the PVA hydrogels to retain their original shape and be extended up to six times their size. This behaviour shows its rubbery and elastic nature and the high mechanical strength [24, 31]. There are various theories proposed in the literature to explain why thermal cycling increases the elastic modulus of PVA. The most accepted theory describes the physical cross-linking process as an entropic reordering phenomena. Water is likely to bind to the polymer by hydrogen bonding. When the solution freezes, ice crystals force the polymer chains close to each other forming high local polymer concentration regions or nuclei. When the material thaws, these nuclei act as crosslinking sites for polymers molecules, which realign and form hydrogen bonds to form crystallites and polymer chain entanglements. The crystalline regions are formed within the polymer rich regions, with further cycling increasing both the size and number of the crystalline regions by repeating the process [19, 32, 27]. On a molecular level, the crystallites of PVA can be described as layered structure, with a double layer of molecules held together by hydroxyl bonds, while weaker van der Waals forces operate between the double layers. This folded chain structure leads to ordered regions (crystallites) within an unordered, amorphous polymer matrix [25]. The mechanical properties of PVA are very unique compared to other polymers. The stress-strain curves for the polymeric materials are initially linear and then curve towards the strain axis. On the other hand, the PVA curve displays an exponential stress-strain curve similar to the characteristics of soft biological tissues, with the curve shifting towards the stress axis.
PVA materials have been reported to be ideal candidates as biomaterials, due to their high degree of swelling, uncomplicated chemical structure, rubbery/elastic nature, non-toxic, non-carcinogenic, and bioadhesive characteristics. Some of the biomedical applications include tissue reconstruction and replacements, cell entrapment and drug delivery, soft contact lens material, wound covering bandage for burn victims, quality control phantom for MR, among other medical applications [32, 25].
Anisotropic Hydrogels
Most research PVA hydrogels has focused on materials exhibiting the normal characteristic of isotropic mechanical behaviour, that is, the mechanical properties of the material are the same regardless of orientation. This is expected due to the random distribution of the polymer chains.
Most tissues, however, including cardiovascular tissues, are composite viscoelastic biomaterials displaying mechanical properties with varying degrees of orientation effects. This orientation effect is due to the organization of the structural protein components such as collagen and elastin within the tissue. This organization gives rise to the unique exponential stress-strain relationship exhibited by soft tissues.
Recently, an anisotropic PVA hydrogel was reported [33] that was able to closely match the stress-strain behaviour of porcine aorta. In this study, it was shown that an anisotropic PVA hydrogel can be produced that displays the exponential response of cardiovascular tissue and also displays the anisotropic behavior of porcine aorta up to 65% strain.
Nanocomposite Hydrogels
A second material system for obtaining improved viscoelastic properties of synthetic and biocompatible replacement tissue materials is that of nanocomposite hydrogels.
Bacterial cellulose has many characteristics that make it valuable for biomedical applications, including its polyfunctionality, hydrophilicity, and biocompatibility [34]. Cellulose is a linear polymer made of glucose molecules linked by β(14) glycosidic linkages. Its chemical formula is (C6H10O5)n. There are four principle sources of cellulose. The majority of cellulose is isolated from plants. A second source is the biosynthesis of cellulose by different microorganisms, including bacteria (acetobacter, aerobacter, pseudomonas), algae, and fungi among others. The other two less common sources include the enzymatic in vitro synthesis starting from cellobiosyl fluoride, and the chemosynthesis from glucose by ring-opening polymerization of benzylated and pivaloylated derivatives [35, 36]. Cellulose is not uniformly crystalline, but ordered regions are extensively distributed throughout the material, and these regions are called crystallites. The long cellulose chains lie side by side held together by hydrogen bonds between the hydroxyl groups. These chains are twisted into structures called microfibrils, which are twisted into fibers [34, 35].
Bacterial cellulose is produced by strains of the bacterium Acetobacter xylinum, which is typically found on decaying fruits, vegetables, vinegar, fruit juices, and alcoholic beverages. It is a Gram-negative, rod shaped and strictly aerobic bacterium. Bacterial cellulose produced has very high purity and contains no lignin, hemicelluloses, pectin, and waxes as plant cellulose does. Therefore, production of bacterial cellulose has the advantage of not requiring the harsh chemical treatment needed for plant cellulose production. This chemical treatment also has the disadvantage of altering the natural structural characteristics of cellulose [34, 35, 36]. Bacterial cellulose differs from plant cellulose with respect to its high crystallinity, ultra-fine network structure, high water absorption capacity, high mechanical strength in the wet state, and availability in an initial wet state [36].
Bacterial cellulose pellicles are formed in static culture. The pellicle has an ultra-fine network structure of ribbons 500 nm wide and 10 nm thick. The ribbons consisted of smaller microfibrils with a width of around 3 nm and a fiber diameter of less than 130 nm compared to the over 14 mm found in birch [35, 36]. Bacterial cellulose including the pellicle possesses a high water retention capacity. Water retention values can reach up to 1000%, which are significantly higher than that for plant cellulose. The water retention is drastically decreased after air-drying the bacterial cellulose and reswelling in water, with values comparable to those of plant cellulose [35, 36].
Bacterial cellulose can also be prepared in shake culture in flasks and in agitated culture in a bioreactor. These approaches are more efficient methods for bacterial cellulose production and are preferred for large scale production of bacterial cellulose.
Bacterial cellulose, being a hydrophilic, highly water swollen and biocompatible natural polymer which is ideally suited to be the reinforcing fibers in the preparation of a nanocomposite material for soft tissue replacement devices. Such nanocomposite material can be created when it is used in combination with PVA.
Uryu [37] reported the formation of a biodegradable polymeric material that can be decomposed in soil. The bacterial cellulose (with ribbon shaped micro-fibrils) that can be biologically decomposed by microbes was mixed with a biodegradable polymeric material to produce an improved composite with higher tensile strength. The bacterial cellulose was produced in a liquid culture medium using different types of, microbes, including Acetobacter xylinum, collected and dried into a powdery state and mixed with the polymer to produce the composite. Various polymers were used, including PVA. The nanocomposites ranged from bacterial cellulose concentrations as low as 1% to 99%. The final composite was dried and used for high-strength cabinets for audio/video apparatus. After the lifetime of the device is reached, the composite material can be buried in the ground for waste disposal and it is eventually decomposed to protect the environment.
U.S. Pat. No. 5,558,861 discloses a hydrogel formed by microbially-produced cellulose that may be complexed with an appropriate auxiliary material (including PVA) for the purposes of reinforcement, change of the specific gravity, immobilization, modification of the affinity, prevention of exudation of the liquid component and the like. This invention teaches a hydrogel that is formed based on the crosslinking of bacterially cellulose, whereby a concentration of PVA can be added for a number of purposes, including reinforcing the crosslinked bacterial cellulose hydrogel.
In contrast to the hydrogel disclosed in U.S. Pat. No. 5,558,861, U.S. patent application Ser. No. 12/216,809 teaches a nanocomposite hydrogel comprising PVA and bacterial cellulose, where the hydrogel is formed by physically crosslinking a PVA solution with a small concentration of bacterial cellulose. This composite hydrogel, in which PVA forms the primary structure of the hydrogel rather than a reinforcing structure, uniquely provides a biocompatible composite hydrogel that exhibits the exponential stress-strain behaviour that is characteristic of many biological tissues.
The nanocomposite hydrogel of U.S. patent application Ser. No. 12/216,809 can be further understood by considering the development of the hydrogel during crosslinking, and the role of the bacterial cellulose in this process. The bacterial cellulose, which forms extensive hydrogen bonds with PVA, is believed to act as a nucleation site for the formation of additional PVA crystallites during crosslinking. Accordingly, the composite hydrogel can be understood to be formed by crosslinking PVA in the presence of bacterial cellulose nanofibers, where the bacterial cellulose promotes additional PVA crystal growth and also contributes to the over strength and compliance properties of the composite [38].
The Need for Additional Stiffness and Anisotropy
For cardiovascular applications, it is important to consider the full strain range that is required in clinical applications. The physiological average strain between diastole and systole for porcine aorta is around 30% strain [33,38,40]. The physiological strain range has also been reported to be between 17 and 49% strain [40, 41]. However, in designing cardiovascular devices, it is necessary to make allowance for higher strain conditions (corresponding to higher systole values) to ensure the material remains elastic at higher strains to ensure durability. Furthermore, a successful replacement must have adequate strength, kink resistance, and must allow sutures to hold under circumferential and axial tension, as well as circumferential and axial compliance [39].
Accordingly, a need remains for a tissue replacement material that can match the highly anisotropic viscoelastic properties of many different types of soft tissues and can also provide improved stiffness beyond typical physiological strain conditions.